Cardiomechanical assessment for cardiac resynchronization therapy

ABSTRACT

A first lead provides therapeutic stimulation to the heart and includes a first mechanical sensor that measures physical contraction and relaxation of the heart. A controller induces delivery of therapeutic stimulation via the first lead. The controller receives signals from the first mechanical sensor indicative of the contraction and relaxation; develops a template signal that corresponds to the contraction and relaxation; and uses the template signal to modify the delivery of therapeutic stimulations. In another arrangement, a second lead, with a second mechanical sensor also provides signals to the controller indicative of contraction and relaxation. The first mechanical sensor is adapted to be positioned at the interventricular septal region of the heart, and the second mechanical sensor is adapted to be positioned in the lateral region of the left ventricle. The controller processes the signals from the first mechanical sensor and the second mechanical sensor to develop a dysynchrony index.

FIELD OF THE INVENTION

The present disclosure relates to an implantable sensor that is capableof measuring longitudinal, radial and torsional strain in the heart. Thestrain data can be used to improve cardiac resynchronization therapytimings for implantable cardiac stimulation devices and systems.

BACKGROUND

Implantable devices for pacing, cardioversion, defibrillation andresynchronization of cardiac electrical and mechanical function arewidely available to prevent and treat arrhythmias and dysynchronousmyocardial mechanics. These disorders can impair cardiac performance byaltering electrical conduction patterns or by changing myocardialcontractility or compliance, both of which result in mechanicaldysfunction.

For example, conduction abnormalities may occur between the atria andthe ventricular chambers. When atrio-ventricular (AV) timing isshortened, ventricular contraction may prematurely terminate the atrialkick produced by the contracting atrium. When AV timing is prolonged,increased ventricular loading from the atria may be lost due toregurgitation during prolonged diastole. Thus, both shortened andprolonged AV timing intervals can affect cardiac output.

Conduction abnormalities between right and left ventricular chambers(inter-ventricular) or within the right or left ventricles(intra-ventricular) can also result in dysynchrony. Dysynchrony occurswhen forces generated in specific regions at inappropriate times causebulging of the chamber walls into adjacent relaxed wall segments, oragainst prematurely closed heart valves. This lack of coordinationduring myocardial contraction may cause a reduction of forward bloodflow and lead to reduced contractile efficiency.

Conduction abnormalities may also result in contractile and complianceabnormalities with cardiac function. For example, conduction delays maycause the left ventricular myocardium to continue to contract even afterthe closure of the aortic valve. This persistent contractile effectcreates post-systolic wall thickening that can reduce left ventriclecompliance and cause a reduction in ventricular end-diastolic volume(pre-load). The reduction in pre-load will reduce stroke volume andcardiac output through the Frank-Starling mechanism. Post-systolic wallthickening and post-systolic myocardial motion are also indicative ofinefficient cardiac effort occurring against a closed aortic valve.

SUMMARY

The present disclosure relates to an implantable cardiomechanical sensorthat is capable of measuring longitudinal, radial and torsionalmotion/deformation (e.g. strain and strain rate) in the heart. In someembodiments, these data can be used to improve cardiac resynchronizationtherapy timings for implantable cardiac stimulation devices and systems.For example, the time to peak strain in the septal and lateral regionsof the myocardium can be compared to determine whether dysynchronyexists. If dysynchrony exists, interventricular timing andatrioventricular timing can be adjusted to reduce the level ofdysynchrony. In other embodiments, the data can be used to detect amyocardial infarction.

In one aspect, the invention relates to an implantable cardiacstimulation device that includes a first lead adapted to be implanted inor on the heart of a patient. The first lead is adapted to providetherapeutic stimulation to the heart of the patient and includes a firstmechanical sensor that obtains measurements indicative of the physicalcontraction and relaxation of the walls of the heart during systole anddiastole. The device also includes a controller that induces a deliveryof therapeutic stimulation to the heart of the patient via the firstlead. The controller receives signals from the first mechanical sensorindicative of the contraction and relaxation of the walls of the heart;develops a template signal that corresponds to the observed contractionand relaxation of the walls of the heart during systole and diastole;and uses the template signal to modify the delivery of therapeuticstimulations being provided to the heart so that the heart's functionduring systole and diastole is improved.

In another aspect, the invention relates to an implantable assessmentdevice that includes a controller configured to accept inputs related tocardiomechanical strain of a lateral region of a left ventricle of aheart and cardiomechanical strain of an interventricular septal regionof the heart. The controller computes an interventricular dysynchronyindex based upon the cardiomechanical strain input from the lateralregion of a left ventricle of a heart. The controller may also determinethe times of peak cardiomechanical strain from the inputs.

In yet another aspect, the invention relates to a method for assessingmyocardial function using cardiomechanical sensors. The method involvesacquiring data over a period of time from a first implanted myocardialmechanical sensor and a second implanted myocardial mechanical sensorseparated by a distance; summating the acquired data from the first andsecond implanted myocardial mechanical sensors; and taking a derivativeof the summated acquired data over the period of time to determine afirst strain rate index.

In yet another aspect, the invention relates to an implantablecardiomechanical assessment system that includes an implantablecardiomechanical sensor system comprising at least a first myocardialmechanical sensor and a second myocardial mechanical sensor. The systemalso includes an implantable controller system coupled to theimplantable cardiomechanical sensor system that is configured to acquiredata over a period of time from the first implanted myocardialmechanical sensor and the second implanted myocardial mechanical sensor,summate the acquired data and calculate a derivative of the summatedacquired data to determine a first strain rate index.

BRIEF DESCRIPTION OF THE DRAWINGS

Further features and advantages of the present disclosure may be morereadily understood by reference to the following description, taken inconjunction with the accompanying drawings, in which:

FIG. 1 is a simplified diagram illustrating an implantable stimulationdevice in electrical communication with at least three leads implantedinto a patient's heart for delivering multi-chamber stimulation andshock therapy.

FIG. 2A is a functional block diagram of a multi-chamber implantablestimulation device illustrating the basic elements of a stimulationdevice which can provide cardioversion, defibrillation and pacingstimulation in four chambers of the heart.

FIG. 2B illustrates a cross section of an embodiment of acardiomechanical electric sensor.

FIG. 3A is a micrograph of a segment of myocardium. FIGS. 3B and 3C areschematic representations of muscle undergoing lengthening andshortening, respectively.

FIG. 4 shows graphs of myocardium tissue velocity and displacementderived from tissue tracking echocardiogram data.

FIG. 5A shows tissue velocity curves for normal heart tissue segmentsfrom the apical region of the heart to the basal region of the heart.

FIG. 5B shows velocity, displacement, strain rate and strain curves formyocardial tissue at the apex, mid-wall and base of the heart.

FIG. 6 is a schematic illustration of two leads, one disposed relativeto the left ventricle and the other disposed relative to the rightventricle, each including a cardiomechanical electric sensor (CMES) forsensing one or more of myocardial motion and deformation.

FIG. 7 is another schematic of the leads of FIG. 6 further depictingfeatures related to myocardial velocity calculation.

FIG. 8A shows two CMESs, one disposed in the interventricular septalregion of the heart and the other in the coronary sinus region of theheart, being used to detect ventricular dysynchrony.

FIG. 8B shows two CMESs, one disposed in the interventricular septalregion of the heart and the other in the lateral region of the leftventricle of the heart, being used to detect ventricular dysynchrony.

FIG. 8C are schematic voltage graphs from the two CMES electrodes.

FIG. 9 depicts a block diagram for determining AV and VV timing settingsthat do not result in dysynchrony by using CMES and the MatrixOptimization Method.

FIG. 10A is a diagram of one CMES disposed in the coronary sinus regionof the heart, being used to detect radial deformation and rotation ofthe heart.

FIG. 10B is a diagram of two CMESs, one disposed in the apical region ofthe heart and the other in the basal region of the heart, being used todetect torsional deformation of the heart.

FIG. 10C is a schematic of two CMES lead configurations, one with CMESmaterial deposited in strips parallel to the lead axis and the otherwith CMES material deposited in a helical configuration around the lead.

FIG. 10D shows a graph of cardiac translation and expansion and a graphof cardiac rotation.

FIG. 11 shows how polarity of the CMES signal can be inferred fromgraphs of an EGM, the raw CMES signal, the longitudinal velocity and therotational velocity.

FIG. 12 is a block diagram for calculating the torsion of the heart.

FIG. 13 shows a surface ECG, a curve representing dCMES/dt and adisplacement curve derived from actual porcine data using an embodimentof a CMES in the right ventricle.

FIG. 14 shows a surface ECG, a curve representing dCMES/dt and adisplacement curve derived from actual porcine data using an embodimentof a CMES in the left ventricle.

FIG. 15 shows a surface ECG, an inverted curve representing dCMES/dt andan inverted displacement curve derived from actual porcine data using anembodiment of a CMES in the left ventricle.

FIG. 16 shows how averaging of the measured mechanical stress waveformscan be synchronized with the detected heart events, such as spontaneousR-waves or stimulated events such as valvular heart sounds.

FIG. 17 is a block diagram for taking CMES measurements when the patientis in a state of relative hypopnea or apnea.

DETAILED DESCRIPTION

Reference will now be made to the drawings wherein like numerals referto like parts throughout. The following description is of the best modepresently contemplated for practicing the invention. This description isnot to be taken in a limiting sense but is made merely for the purposeof describing the general principles of the invention. The scope of theinvention should be ascertained with reference to the issued claims. Inthe description of the invention that follows, like numerals orreference designators will be used to refer to like parts or elementsthroughout.

In one embodiment, as shown in FIG. 1, an implantable cardiacstimulation device 10 is in electrical communication with a patient'sheart 12 by way of three leads, 20, 24 and 30, suitable for deliveringmulti-chamber stimulation and shock therapy. To sense atrial cardiacsignals and to provide right atrial chamber stimulation therapy, thestimulation device 10 is coupled to an implantable right atrial lead 20having at least an atrial tip electrode 22, which typically is implantedin the patient's right atrial appendage.

To sense left atrial and ventricular cardiac signals and to provide leftchamber pacing therapy, the stimulation device 10 is coupled to a“coronary sinus” lead 24 designed for placement in the “coronary sinusregion” via the coronary sinus ostium (OS) for positioning a distalelectrode adjacent to the left ventricle and/or additional electrode(s)adjacent to the left atrium. As used herein, the phrase “coronary sinusregion” refers to the vasculature of the left ventricle, including anyportion of the coronary sinus, great cardiac vein, left marginal vein,left posterior ventricular vein, middle cardiac vein, and/or smallcardiac vein or any other cardiac vein accessible by the coronary sinus.

Accordingly, an exemplary coronary sinus lead 24 is designed to receiveatrial and ventricular cardiac signals and to deliver left ventricularpacing therapy using at least a left ventricular tip electrode 26, leftatrial pacing therapy using at least a left atrial ring electrode 27,and shocking therapy using at least a left atrial coil electrode 28.

The stimulation device 10 is also shown in electrical communication withthe patient's heart 12 by way of an implantable right ventricular lead30 having, in this embodiment, a right ventricular tip electrode 32, aright ventricular ring electrode 34, a right ventricular (RV) coilelectrode 36, and a superior vena cava (SVC) coil electrode 38.Typically, the right ventricular lead 30 is transvenously inserted intothe heart 12 so as to place the right ventricular tip electrode 32 inthe right ventricular apex so that the RV coil electrode will bepositioned in the right ventricle and the SVC coil electrode 38 will bepositioned in the superior vena cava. Accordingly, the right ventricularlead 30 is capable of receiving cardiac signals, and deliveringstimulation in the form of pacing and shock therapy to the rightventricle.

The right atrial lead 20, the coronary sinus lead 24, and the rightventricular lead 30 can all incorporate cardiomechanical electric sensor(CMES) material so that the leads can function to provide cardiacmechanical motion data as described herein.

As illustrated in FIG. 2A, a simplified block diagram is shown of themulti-chamber implantable stimulation device 10, which is capable oftreating both fast and slow arrhythmias with stimulation therapy,including cardioversion, defibrillation, and pacing stimulation. While aparticular multi-chamber device is shown, this is for illustrationpurposes only and one of skill in the art could readily duplicate,eliminate or disable the appropriate circuitry in any desiredcombination to provide a device capable of treating the appropriatechamber(s) with cardioversion, defibrillation and pacing stimulation.

The housing 40 for the stimulation device 10, shown schematically inFIG. 2A, is often referred to as the “can”, “case” or “case electrode”and may be programmably selected to act as the return electrode for allpacemaker “unipolar” modes. The housing 40 may further be used as areturn electrode alone or in combination with one or more of the coilelectrodes, 28, 36 and 38, for shocking purposes. The housing 40 furtherincludes a connector (not shown) having a plurality of terminals 42, 44,46, 48, 52, 54, 56, and 58 (shown schematically and, for convenience,the names of the electrodes to which they are connected are shown nextto the terminals). As such, to achieve right atrial sensing and pacing,the connector includes at least a right atrial tip terminal (A_(R) TIP)42 adapted for connection to the atrial tip electrode 22.

To achieve left chamber sensing, pacing and shocking, the connectorincludes at least a left ventricular tip terminal (V_(L) TIP) 44, a leftatrial ring terminal (A_(L) RING) 46, and a left atrial shockingterminal (A_(L) COIL) 48, which are adapted for connection to the leftventricular tip electrode 26, the left atrial ring electrode 27, and theleft atrial coil electrode 28, respectively.

To support right chamber sensing, pacing and shocking, the connectorfurther includes a right ventricular tip terminal (V_(R) TIP) 52, aright ventricular ring terminal (V_(R) RING) 54, a right ventricularshocking terminal (RV COIL) 56, and an SVC shocking terminal (SVC COIL)58, which are adapted for connection to the right ventricular tipelectrode 32, right ventricular ring electrode 34, the RV coil electrode36, and the SVC coil electrode 38, respectively.

At the core of the stimulation device 10 is a programmablemicrocontroller 60 which controls the various modes of stimulationtherapy. As is well known in the art, the microcontroller 60 typicallyincludes a microprocessor, or equivalent control circuitry, designedspecifically for controlling the delivery of stimulation therapy and mayfurther include RAM or ROM memory, logic and timing circuitry, statemachine circuitry, and I/O circuitry. Typically, the microcontroller 60includes the ability to process or monitor input signals (data) ascontrolled by a program code stored in a designated block of memory. Thedetails of the design and operation of the microcontroller 60 are notcritical to the invention. Rather, any suitable microcontroller 60 maybe used that carries out the functions described herein. The use ofmicroprocessor-based control circuits for performing timing and dataanalysis functions are well known in the art.

As shown in FIG. 2A, an atrial pulse generator 70 and a ventricularpulse generator 72 generate pacing stimulation pulses for delivery bythe right atrial lead 20, the right ventricular lead 30, and/or thecoronary sinus lead 24 via an electrode configuration switch 74. It isunderstood that in order to provide stimulation therapy in each of thefour chambers of the heart, the atrial and ventricular pulse generators70, 72 may include dedicated, independent pulse generators, multiplexedpulse generators, or shared pulse generators. The pulse generators 70,72 are controlled by the microcontroller 60 via appropriate controlsignals, 76 and 78, respectively, to trigger or inhibit the stimulationpulses.

The microcontroller 60 further includes timing control circuitry 79which is used to control the timing of such stimulation pulses (e.g.,pacing rate, atrio-ventricular (AV) delay, atrial interconduction (A-A)delay, or ventricular interconduction (V-V) delay, etc.) as well as tokeep track of the timing of refractory periods, PVARP intervals, noisedetection windows, evoked response windows, alert intervals, markerchannel timing, etc., which is well known in the art.

The switch 74 includes a plurality of switches for connecting thedesired electrodes to the appropriate I/O circuits, thereby providingcomplete electrode programmability. Accordingly, the switch 74, inresponse to a control signal 80 from the microcontroller 60, determinesthe polarity of the stimulation pulses (e.g., unipolar, bipolar,combipolar, etc.) by selectively closing the appropriate combination ofswitches (not shown) as is known in the art. In this embodiment, theswitch 74 also supports simultaneous high resolution impedancemeasurements, such as between the case or housing 40, the right atrialelectrode 22, and right ventricular electrodes 32, 34 as described ingreater detail below.

Atrial sensing circuits 82 and ventricular sensing circuits 84 may alsobe selectively coupled to the right atrial lead 20, coronary sinus lead24, and the right ventricular lead 30, through the switch 74 fordetecting the presence of cardiac activity in each of the four chambersof the heart. Accordingly, the atrial (ATR. SENSE) and ventricular (VTR.SENSE) sensing circuits 82, 84 may include dedicated sense amplifiers,multiplexed amplifiers, or shared amplifiers. The switch 74 determinesthe “sensing polarity” of the cardiac signal by selectively closing theappropriate switches, as is also known in the art. In this way, theclinician may program the sensing polarity independently of thestimulation polarity.

Each sensing circuit 82, 84 preferably employs one or more low power,precision amplifiers with programmable gain and/or automatic gaincontrol, bandpass filtering, and a threshold detection circuit, as knownin the art, to selectively sense the cardiac signal of interest. Theautomatic gain control enables the device 10 to deal effectively withthe difficult problem of sensing the low amplitude signalcharacteristics of atrial or ventricular fibrillation. The outputs ofthe atrial and ventricular sensing circuits 82, 84 are connected to themicrocontroller 60 which, in turn, are able to trigger or inhibit theatrial and ventricular pulse generators 70, 72 respectively, in a demandfashion in response to the absence or presence of cardiac activity inthe appropriate chambers of the heart.

For arrhythmia detection, the device 10 utilizes the atrial andventricular sensing circuits 82, 84 to sense cardiac signals todetermine whether a rhythm is physiologic or pathologic. As used herein“sensing” is reserved for the noting of an electrical signal, and“detection” is the processing of these sensed signals and noting thepresence of an arrhythmia. The timing intervals between sensed events(e.g., P-waves, R-waves, and depolarization signals associated withfibrillation) are then classified by the microcontroller 60 by comparingthem to a predefined rate zone limit (i.e., bradycardia, normal, lowrate VT, high rate VT, and fibrillation rate zones) and various othercharacteristics (e.g., sudden onset, stability, physiologic sensors, andmorphology, etc.) in order to determine the type of remedial therapythat is needed (e.g., bradycardia pacing, anti-tachycardia pacing,cardioversion shocks or defibrillation shocks, collectively referred toas “tiered therapy”).

Cardiac signals are also applied to the inputs of an analog-to-digital(A/D) data acquisition system 90. The data acquisition system 90 isconfigured to acquire intracardiac electrogram (IEGM) signals, convertthe raw analog data into a digital signal, and store the digital signalsfor later processing and/or telemetric transmission to an externaldevice 102. The data acquisition system 90 is coupled to the rightatrial lead 20, the coronary sinus lead 24, and the right ventricularlead 30 through the switch 74 to sample cardiac signals across any pairof desired electrodes.

The microcontroller 60 is further coupled to a memory 94 by a suitabledata/address bus 96, wherein the programmable operating parameters usedby the microcontroller are stored and modified, as required, in order tocustomize the operation of the stimulation device 10 to suit the needsof a particular patient. Such operating parameters define, for example,pacing pulse amplitude, pulse duration, electrode polarity, rate,sensitivity, automatic features, arrhythmia detection criteria, and theamplitude, waveshape and vector of each shocking pulse to be deliveredto the patient's heart 12 within each respective tier of therapy.

Advantageously, the operating parameters of the implantable device 10may be non-invasively programmed into the memory 94 through a telemetrycircuit 100 in telemetric communication with the external device 102,such as a programmer, transtelephonic transceiver, or a diagnosticsystem analyzer. The telemetry circuit 100 is activated by themicrocontroller by a control signal 106. The telemetry circuit 100advantageously allows IEGMs and status information relating to theoperation of the device 10 (as contained in the microcontroller 60 ormemory 94) to be sent to the external device 102 through an establishedcommunication link 104.

In the preferred embodiment, the stimulation device 10 further includesa physiologic sensor 108, commonly referred to as a “rate-responsive”sensor because it is typically used to adjust pacing stimulation rateaccording to the exercise state of the patient. However, thephysiological sensor 108 may further be used to detect changes incardiac output, changes in the physiological condition of the heart, ordiurnal changes in activity (e.g., detecting sleep and wake states).Accordingly, the microcontroller 60 responds by adjusting the variouspacing parameters (such as rate, AV Delay, V-V Delay, etc.) at which theatrial and ventricular pulse generators 70, 72 generate stimulationpulses.

The stimulation device additionally includes a battery 110 whichprovides operating power to all of the circuits shown in FIG. 2A. Forthe stimulation device 10, which employs shocking therapy, the battery110 must be capable of operating at low current drains for long periodsof time and then be capable of providing high-current pulses (forcapacitor charging) when the patient requires a shock pulse. The battery110 must also have a predictable discharge characteristic so thatelective replacement time can be detected. Accordingly, embodiments ofthe device 10 including shocking capability preferably employlithium/silver vanadium oxide batteries. For embodiments of the device10 not including shocking capability, the battery 110 will preferably belithium iodide or carbon monoflouride or a hybrid of the two.

As further shown in FIG. 2A, the device 10 is shown as having animpedance measuring circuit 112 which is enabled by the microcontroller60 via a control signal 114.

In the case where the stimulation device 10 is intended to operate as animplantable cardioverter/defibrillator (ICD) device, it must detect theoccurrence of an arrhythmia, and automatically apply an appropriateelectrical shock therapy to the heart aimed at terminating the detectedarrhythmia. To this end, the microcontroller 60 further controls ashocking circuit 116 by way of a control signal 118. The shockingcircuit 116 generates shocking pulses of low (up to 0.5 joules),moderate (0.5-10 joules), or high energy (11 to 40 joules), ascontrolled by the microcontroller 60. Such shocking pulses are appliedto the patient's heart 12 through at least two shocking electrodes, andas shown in this embodiment, selected from the LA coil electrode 28, theRV coil electrode 36, and/or the SVC coil electrode 38. As noted above,the housing 40 may act as an active electrode in combination with the RVelectrode 36, or as part of a split electrical vector using the SVC coilelectrode 38 or the LA coil electrode 28 (i.e., using the RV electrodeas a common electrode).

Cardioversion shocks are generally considered to be of low to moderateenergy level (so as to minimize pain felt by the patient), and/orsynchronized with an R-wave and/or pertaining to the treatment oftachycardia. Defibrillation shocks are generally of moderate to highenergy level (i.e., corresponding to thresholds in the range of 5-40joules), delivered asynchronously (since R-waves may be toodisorganized), and pertaining exclusively to the treatment offibrillation. Accordingly, the microcontroller 60 is capable ofcontrolling the synchronous or asynchronous delivery of the shockingpulses.

A variety of diseases such as cardiomyopathy, congestive heart failure,hypertrophic cardiomyopathy, aortic stenosis and ischemic heart diseaseshow characteristic abnormalities in myocardial strain, myocardialtissue velocity and myocardial tissue displacement, rotation andtorsion. Tissue Doppler imaging (TDI) data is used to derive myocardialstrain and strain rate by analysis of regional disparities in tissuevelocity or the spatial location of ultrasonic reflectors (speckletracking) as a function of time. This information is used clinically toevaluate properties of myocardial motion and deformation that provideinsight into the electromechanics of the heart.

In some embodiments, lead-based sensors may be used as an alternative toTDI for generating quantitative information which relates to the sameproperties such as myocardial strain, myocardial strain rate, myocardialtissue velocity and myocardial tissue displacement, rotation andtorsion. Sensors capable of acquiring this data can be used formonitoring purposes and communicate information related to cardiacperformance and dysynchrony to the clinician. The same data can be usedas part of a closed loop system for CRT timing.

Piezoelectric materials will generate a voltage when subject tomechanical stress or strain, with the magnitude of voltage dependentupon the magnitude of the stress or strain. In some embodiments, sensorscomprised of piezoelectric material and positioned in locations optimalfor detection of cardiac deformation and/or motion generate raw signalsof cardiac mechanical data that can be further processed into myocardialstrain, myocardial strain rate, myocardial tissue velocity andmyocardial tissue displacement, rotation and torsion data.

Embodiments of CMESs may comprise one or more piezoelectric transducers,which convert mechanical motion into electrical signals. As illustratedin cross section in FIG. 2B, in some embodiments, a CMES 200 comprises atubular and/or annular piezoelectric element 210, either self-supportingor disposed on a supporting structure. In some embodiments, conductors220 contact the inner and outer surfaces of the tubular or annularelement 210. Electrical connections 230 are coupled to the conductors220.

In preferred embodiments, the sensor 200 is dimensioned forincorporation into a lead. For example, in some embodiments, the outerdiameter of the sensor 200 is similar to the outer diameter of a lead,permitting the sensor to be disposed at any position along a leadwithout causing a profile change that could affect placement of thelead. In some embodiments, one or more of an electrode and/or othersensors is disposed over at least a portion of the sensor 200. Alongitudinal passageway 240 through the sensor 200 in the illustratedembodiment permits routing electrical and/or other types of connectionstherethrough, for example, from one or more electrodes and/or sensorsdisposed on the same lead.

The conductors 220 comprise any suitable material known in the art, forexample, titanium, titanium alloy, titanium nitride, platinum, platinumalloy, carbon, niobium, niobium alloy, tantalum, tantalum alloy, gold,combinations, and the like. In some embodiments, a patient's tissue isused as one of the conductors. In some embodiments, an elastomer isdisposed over the sensor 200 (not illustrated). Preferred elastomers arebiocompatible, including, for example, silicones, polyurethanes,ethylene-propylene copolymers, fluorinated elastomers, combinations, andthe like.

In some embodiments, the piezoelectric element comprises a relativelyhard material, thereby permitting reliable measurements with only smalldeflections of the piezoelectric material. Preferred piezoelectricmaterials are biocompatible, for example, ceramic piezoelectricmaterials, including ceramic ferroelectric particles, lead zirconatetitanate (lead zirconium titanate, PZT), barium titanate, sodiumpotassium niobate, and the like. In some embodiments, the piezoelectricmaterial comprises Na_(0.5)K_(0.5)NbO₃, for example, as described inU.S. Pat. No. 6,526,984. Other piezoelectric materials ordeformation-based sensors may also be used.

One preferred sensor configuration comprises a piezoelectric materialthat is thin and covers a large amount of myocardial tissue surfacearea. Covering a large surface area provides global deformation data incomparison to the local information acquired by CMES material depositedin a smaller region. In order for data to be representative ofmyocardial deformation the CMES preferably contacts myocardium, andthus, the CMES is preferably located along the distal portion of a leadbody and contours along either a large caliber coronary sinus lead orthe epicardial surface if the CMES is deployed via a limited thoracotomy(e.g., a pericardial or epicardial approach).

In other embodiments, the CMES comprises a conductive polymer that has aresistance that changes as a function of strain. By measuring theresistance of the conductive polymer, the strain can be determined. Theconductive polymer can be polyacetylene, polyaniline, polypyrrole or anyother suitable conductive polymer.

In some embodiments that use piezoelectric materials, the raw CMESsignal is a measurement of deformation (strain), and can be expressed inunits of voltage. Referring to FIG. 3A, which depicts a micrograph ofsome isolated cardiac muscle fibers 304 from the heart, contraction andrelaxation of the myocytes may be quantified by the deformation ofadjacent mechanical sensors. Strain in the myocardium may be measured bythe change in relevant length of myocardium:

Strain=e=(L−Lo)/Lo   (Eq. 1)

The strain (e) given by Eq. 1 is a dimensionless quantity. Strain ismeasure of a fractional change from unstressed dimension given by theunstressed zero length. Referring to FIGS. 3B and 3C, an expansion tothe muscle fiber 304 length L 302 from the initial length Lo 300represents a positive strain, while a compression and dimensionalshortening represents a negative strain L 302.

A first order derivative of the raw strain signal with respect to timegenerates a measure of the deformation (strain) rate. The calculatedquantity, strain rate, with the unit 1/s is a measure of the rate ofdeformation and is equivalent to the shortening or lengthening velocityper fiber length.

The microcontroller 60 can also comprise circuitry to process dataobtained by the CMES as described herein as part of a closed loopsystem. Alternatively, the data obtained by the CMES can be communicatedto an external device 102 and processed thereafter.

Derivation of CMES Derived Deformation and Velocity Indices

FIG. 4 depicts a velocity curve 400 of a region of the myocardium 402generated by tissue tracking data derived from echocardiographs 404.Tissue tracking images are two-dimensional maps that display color-codedtissue velocity information and can be used to identify wall motionabnormalities and to estimate regional strain or shortening of themyocardium. Tissue tracking may be particularly useful in identifyingwall motion abnormalities that may be treated with resynchronizationtherapy or may be used to optimize resynchronization therapy. A time ofintegral of the velocity curve 400 yields a displacement curve 406 ofthe same region of myocardium.

FIG. 5A depicts echo-generated tissue velocity curves 400 in the apicalto basal regions of the heart. The y-axis 502 represents the velocity,the x-axis 504 represents time and the area under the curve 506, whichcan be obtained by integrating the velocity curve 400, represents tissuedisplacement. FIG. 5B shows a velocity curve 400, a displacement curve406 which can be obtained by integrating the velocity curve 400, astrain rate curve 510 and a strain curve 512 which can be obtained byintegrating the strain rate curve 510 in the apex 514, mid-wall 516 andbase 518 of the heart.

As shown in FIGS. 5A, and 5B, along a longitudinal axis generallyparallel to the spine, the heart contracts and moves from base to apexduring systole. The heart relaxes and moves in the opposite direction,from apex to base, during diastole. The basal regions generally move agreater distance, an average of approximately 12 mm at most basalsegments, than the apical regions, which move approximately 0-2 mm atcardiac apex. A measurement of the relative difference in distance thatany two regions traverse will generate longitudinal deformation (strain)information. Echocardiographic techniques such as tissue trackingdemonstrate this displacement phenomenon as well as characteristics ofvelocity, strain and strain rate.

In some embodiments, as shown in FIG. 6, a CMES-bearing device 600, suchas an LV lead, is positioned with respect to the left ventricle (LV)such that a portion of the lead incorporating CMES material 604 issubstantially parallel to the cardiac central longitudinal axis (CLA) tothereby acquire longitudinal deformation information. Similarinformation can be acquired by positioning another CMES-bearing device602, such as an RV lead, such that a portion of the lead incorporatingCMES material 606 is substantially parallel with the CLA. Both CMESdevices 600, 602 generate data related to the motion of the cardiac apexrelative to the base if the CMES material 604, 606 covers enough surfacearea along the CLA of the heart and contacts myocardial tissue. Thegreater the distance (base to apex) the CMES material 604, 606traverses, the greater the amount of resultant deformation and the moreglobal the representation of cardiac motion will be. In the case of theRV lead 602, additional CMESs 608, 610 may be included. Basally locatedCMES 610 will deform more than apically located CMES 608 and thus bemore sensitive to changes in global cardiac geometry.

In some embodiments, regional contractile information can be generatedfrom CMES material that covers a short distance. In normal hearts orhearts with global decreases in contractility (strain, deformation) sucha reduced surface area electrode can provide information about globalcardiac contractile function because any regional properties arehomogeneous with global properties (e.g. dilated cardiomyopathy).However, in more anisotropic conditions, whether in the space domain ortime domain, such as ischemic cardiomyopathy or electromechanicaldysynchrony, respectively, regional information provides littleinformation about global cardiac contractile function. As the heart isembryologically and structurally derived from a single muscle band thathas certain deformation properties, tethering effects (e.g. regionalmyocardial shortening has a pulling effect on surrounding myocardium)create some degree of interrelation between regional and global cardiacdeformation. Thus, CMES acquired data in the longitudinal axis willprovide clinically relevant information if the material covers enoughsurface area (e.g. longitudinal lead length).

Relative differences in tissue velocity can be used to determinemyocardial strain rate derived by using the strain rate equation. Thistechnique is implemented in sophisticated echocardiography machines thatare capable of tissue Doppler imaging for quantifying regionalmyocardial strain rate, strain, velocity and displacement. This equationcan be similarly applied herein to derive analogous indices descriptiveof the same myocardial properties using implanted CMES technology. Thestrain rate (SR) equation is:

SR=(Vb−Va)/x   (Eq. 2)

where Vb and Va represent regional velocities at points b and a,respectively, SR=strain rate and x=length between points a and b. Thecalculated strain rate is representative of the myocardial deformationin the region encompassing points a and b where the tissue velocitieswere measured. Similarly, Eq. 2 can be utilized to derive estimatedtissue velocity information of cardiac motion by using the strain ratebetween points a and b measured with a CMES sensor capable of measuringstrain. Taking the derivative of the strain with respect to time yieldsthe strain rate, which can then be used in Eq. 2 to determine velocityinformation.

For example in some embodiments, as shown in FIG. 7, if CMES material604 in series and in contact with the myocardium at points a 700 and b702, which are separated by distance x 704, the resultant summeddeformation voltage Vsum that is generated by the series sensor material604 provides strain information produced along distance x 704 and can beused to derive a deformation index. This property may also be acquiredby depositing CMES material 604 along a relatively long portion of animplanted lead in contact with myocardium. In some embodiments, distancex 704 is in a range of about 4 mm to about 30 mm, preferably about 5 mmto about 20 mm, and most preferably about 5 mm to about 10 mm.

The first derivative of the signal generated from CMES deformationbetween points a and b as a function of time, dVsum/dt=dCMES/dt, isproportionate to SR and can be used to derive a SR index that can beplotted as a function of time. The integration of the SR index can beperformed to derive an index of strain, which in some embodiments is anindex of longitudinal strain. The measure of strain or strain ratebetween points a and b can be used to detect a myocardial infarct bycomparing the measured strain or strain rate values with expected ornormal strain or strain rate values. Abnormally low strain or strainrate values may indicate the presence of a myocardial infarct.

dCMES/dt=SR Index   (Eq. 3)

In order to derive regional velocity information, a velocity index, Vi,can be defined that is representative of the spatial velocity gradientbetween points a and b, having a distance x, where Vb and Va representregional velocities at points b and a, respectively. Rearranging Eq. 2,the strain rate equation, and substituting Vi for Vb−Va yields:

Vi=Vb−Va=(SR Index)*(x)   (Eq. 4)

Thus, by combining Eq. 3 and Eq. 4, the CMES derived Velocity Index, Vi,equals the first order derivative, d(CMES)/dt, multiplied by x, where xis the span of the distance between CMES electrodes a and b (or lengthalong a lengthy CMES electrode). This index can be expressed in units,Voltage-cm/sec.

CMES Velocity Index=d(CMES)/dt*x   (Eq. 5)

This index can be measured instantaneously by using d(CMES)/dt max ormeasured as a function of time during the cardiac cycle. This indexgenerally parallels Tissue Doppler measurements of myocardial velocity.Integration of this velocity waveform will provide displacementinformation and measurements such as peak longitudinal displacement canbe derived.

An alternate means of deriving an index of myocardial velocity is bydefining the pure CMES signal as a measurement of motion (e.g. velocity,acceleration). In order for the CMES to represent motion rather thandeformation, the CMES is preferably not fixated to myocardium and isinstead relatively free floating.

Dysynchrony Index

In some embodiments, as shown in FIGS. 8A and 8B, if two or more CMESs800, 802 are deployed in interventricular septal and LV lateral regions,respectively, information about dysynchrony can be derived. One CMES 802can be deployed in the LV lateral region via the coronary sinus region(FIG. 8A) or by a transeptal approach (FIG. 8B). Alternatively, apericardial approach (not shown) may be used to place the CMES 802 inthe LV lateral region. Though electromechanical dysynchrony is ananisotropic property, differences between septal and lateral wall motionare often seen in patients suffering from dysynchrony and suchmeasurements are considered specific indicators of patients who respondto cardiac resynchronization therapy (CRT) device implants. Thus, asshown in FIG. 8C, time of peak CMES Voltage (tCMESpeak) from the septalsensor 800 and lateral sensor 802 located in the distal portion of leadand proximate to myocardium may be used together to provide a CMESDysynchrony Index or other parameter that is indicative of the time topeak myocardial strain, which is a currently utilized ultrasonicmeasurement of dysynchrony. In another embodiment, RV apically placedleads may generate similar information if the CMES material deformationis congruent with septal deformation during the cardiac cycle.

$\begin{matrix}{{{CMES}\mspace{14mu} {Dysynchrony}\mspace{14mu} {Index}} = \frac{\left( {{tCMES}_{peak}\mspace{14mu} {septal}} \right)}{\left( {{tCMES}_{peak}\mspace{14mu} {lateral}} \right)}} & \left( {{Eq}.\mspace{11mu} 6} \right)\end{matrix}$

Alternatively, time to peak d(CMES)/dt, which will parallel measurementsof time to peak SR, can be used instead to calculate the CMESDysynchrony Index.

Other features of the CMES signal can be used for timing (e.g. time ofonset of CMES voltage waveform (Vcmes) or time to peak dVcmes/dt).Generally, the relative timings of the CMES generated signals inopposing regions of interest, for example myocardial wall segments, canbe utilized for deriving a dysynchrony index.

As the CMES Dysynchrony Index approaches a value of one, conditions ofsynchrony will be present. Ideally, this time will occur during thelatter portion of the systolic ejection phase, when strain 512 ismaximal in normal hearts, as shown in FIG. 5B. As changes in intervaltiming occur, the index may be followed and the programmed intervalsthat yield an index that is closest to unity will be optimal. Changes inatrioventricular (AV) and interventricular (VV) timing can be made suchthat multiple permutations of AV and VV intervals are evaluated becausechanges in AV timing and VV timing do not have mutually exclusiveeffects on cardiac synchrony or systolic or diastolic performance. Asshown in FIG. 9, an array or matrix 900 of several AV and VV intervalscan be tested using a Matrix Optimization Method (MOM) while the CMESDysynchrony Index (CMES DI) is evaluated for each permutation. MOM isdescribed in greater detail in U.S. Pat. No. 7,010,347, hereinincorporated by reference in its entirety. Regarding CMES DI evaluation,at block 904 CMES DI is calculated for a set of current AV and VVintervals. At block 906 the calculated CMES DI is compared to unity plusor minus a default value (e.g. a programmable standard deviation). Ifthe CMES DI does not equal unity plus or minus the default value, theprocess returns to block 902 where another set of AV and VV intervalsare selected and block 904 where another CMES DI is calculated. OnceCMES DI equals unity plus or minus a default value, the tested AV and VVintervals are programmed into the device at block 912. The standarddeviation can be derived by analysis of previous values during earlieroptimization efforts.

The CMES Dysynchrony Index may also be used with intracardiacelectrogram (IEGM) data for monitoring electromechanical dysynchrony inthe heart. If electromechanical dysynchrony is detected, lead based CMESelectrodes, as described herein, can be used to implementresynchronization timing therapy as part of a closed loop system. See,for example, U.S. Pat. No. 7,010,347, previously incorporated byreference.

Radial Deformation and Cardiac Rotation

With reference to FIGS. 10A and 10B, in some embodiments, CMESs 1010 maybe deployed circumferentially along the proximal to lateral portion ofthe main coronary sinus branch (endovascular leads) (FIG. 10A), or alongthe AV groove (pericardial leads) (FIG. 10B). In these arrangements,parameters of radial deformation and motion can be derived. Radialstrain can be used as a global cardiac performance index. However,radial strain is subject to regional effects and the performance of moreapical segments may not be well represented, leading to the possibilitythat regional pathology (e.g. an mid-cavitary or apical infarct) willnot be detected.

In some embodiments, a lead configuration where the CMES is in closeproximity to tissue and not free-floating may be utilized to deriverotational velocity information using Eq. 5, thereby providing an indexof basal cardiac rotational velocity. If this data is also acquiredabout the cardiac apex, which is preferably obtained with a pericardialor epicardial lead deployed using a sub-xyphoid approach as shown inFIG. 10B, relative rotational data can be acquired for derivation of atorsion index. In normal hearts, the cardiac base rotates in an oppositedirection from the apex. For example, during isovolumic contraction, thebase rotates counter-clockwise while the apex rotates clockwise. Theopposite motion occurs during isovolumic relaxation as shown in FIG.10D. In FIG. 10D, curve 1000 represents tissue velocity as a function oftime for basilar systolic counter-clockwise rotation and diastoliccounter-clockwise rotation. Curve 1002 is apical systolic clockwiserotation and diastolic counter-clockwise rotation. This torsion effectis pivotal in generating forces that contribute to isovolumiccontraction, aortic valve opening and systolic forward flow and adiastolic suction effect that contributes to early diastolic rapidfilling during isovolumic relaxation. Time T 1004 is diastolic fillingtime where no torsion is present and the heart translates and expands1006 rather than rotates. Identification of this timeframe usingintracardiac electrograms (e.g. just before and after the P wave) canassist in temporal labeling of the generated CMES signals (see below).Leads placed using a pericardial or epicardial approach are generallymore appropriately oriented for generation of clinically relevant CMESsignals.

In some embodiments, circumferential deformation effects (i.e. systoliccircumferential shortening) will contribute to the raw radial CMESsignal data. Thus, the derived rotational velocity information includesboth the actual rotational velocity information plus a contribution fromcircumferential deformation effects. In studies using Tissue VelocityImaging, the estimated amount of contribution of circumferentialdeformation to the measured velocity data is approximately 13% in normalpatients and under 5% in patients with Class III or IV heart failure andejection fraction less than 40% (personal, unpublished data). Thus,application of Eq. 5 to radially derived CMES data will provide arelatively accurate index of pure cardiac rotational velocity with somecontribution from the effects of circumferential deformation. The amountof contribution of circumferential deformation and of rotationalvelocity to the data acquired will also relate to the amount of contactthe sensor has with underlying tissue. Nonetheless, this cardiacperformance index is a useful blend of rotational velocity andcircumferential contractile properties. If directional information canbe derived (e.g. clockwise vs. counter-clockwise) from sensors 1010 and1012 positioned in the apex and base, respectively, as shown in FIG.10B, a torsion index can be obtained by adding the measured indices (inessence adding the absolute values as the rotational vectors areopposite). Patients with more advanced heart failure will have lessrotation and/or torsion and less circumferential deformation with aresultant translational motion without significant rotationalvelocities. Thus, the resulting CMES rotational and torsional index willbe less in these patients. Integration of rotational or torsionaldCMES/dt will derive a rotational or torsional displacement index,respectively.

Inferred Polarity

In some embodiments, embedding CMES material on an implantable lead suchthat the voltage generated relies on the direction of deformation willallow the derivation of more accurate representations of actualphysiologic properties. For example, as shown in FIG. 10C, the CMESmaterial 1014 can be placed in strips parallel 1016 to the long axis ofa lead body 1018 or in a helical fashion 1020 about the lead body 1018.If a lead placed about the AV ring (basal location) has CMES materialembedded parallel to the lead long axis, the raw voltage signalgenerated is more of a function of radial deformation. If the CMESmaterial runs in a helical fashion about the long axis of the lead, theraw voltage signal generated is more a function of circumferentialdeformation.

In some embodiments, if basal and apical CMES electrodes 1010 and 1012are designed to derive rotational indices as shown in FIG. 10B, certainassumptions about direction of deformation may be made. For example, ifdeformation of the CMES material causes it to expand, a differentvoltage waveform will be generated than if the CMES material contracts.The waveform polarity will not be significantly different as the cardiacforces causing the deformation from the original length result in avoltage signal regardless of material contraction or expansion. Certaincharacteristics of the raw voltage signal (e.g. relative positive tonegative polarity in a signal that is not rectified), however, will beseen as a result of CMES material contraction rather than expansion andvice versa. Signal processing can be applied to derive such polarityinformation.

Referring to FIG. 11, the timing of the voltage signal 1100 will relateto systolic (e.g. isovolumic contraction) and diastolic (e.g. isovolumicrelaxation) deformation voltages, Vsys and Vdias, respectively.Isovolumic contraction (IVC) causes a steep rise in longitudinalmyocardial tissue velocity 1102, rotational velocity 1104, longitudinaland radial deformation (strain). IVC typically occurs shortly afterdepolarization. Vsys will typically occur shortly after theelectrocardiogram 1106 (EGM) R wave 1108, while Vdias will typicallyoccur thereafter, before the EGM P wave 1110. In a normal heart,deformation of a lead-based CMES with specific lead orientation andmaterial characteristics (e.g. parallel to lead body, parallel to thecardiac CLA) can be expected to be a result of longitudinal rather thanradial deformation. Similarly, inferences about whether a voltage signalis generated as a result of longitudinal systolic contraction ratherthan diastolic expansion may be made. For example, systolic longitudinalcontraction will occur during IVC. This will lead to contraction of CMESmaterial positioned along the length of a lead that is parallel to thecardiac CLA. The resulting waveform will occur after the EGM R wave andthus, the second CMES voltage waveform, Vdias, can be inferred to be aresult of material expansion during diastole. Likewise, the EGM R toVsys interval will be shorter than the EGM R to Vdias interval, and theinterval from Vdias to the next EGM R will be shorter than the intervalfrom Vsys to the next EGM R. Furthermore, Vdias will often be of loweramplitude than Vsys as the forces generated from diastolic expansion(isovolumic relaxation, IVR) are of less amplitude and are generatedmore slowly than systolic contraction that occurs during IVC.Identification of the temporal relationship of these waveforms to theintracardiac P wave will assist in labeling a given signal as onegenerated from contraction and not relaxation. These temporal andmorphologic signal characteristics will allow the system to inferpolarity of deformation information. Apically located CMES sensors 1112will have an assigned polarity that is different than basally locatedCMES sensors 1114 which is represented on bottom of FIG. 11 as Aa andAb, respectively. Under normal circumstances these deflections (withinferred polarity) will be in the opposite direction as shown, though inpatients with congestive heart failure the amplitude will be less andthe direction of these signals may be similar (secondary to translationwithout rotation and impaired circumferential shortening).

In the pathologic heart, these temporal and morphologic signalcharacteristics may be less accurate and signal processing fordetermination of inferred polarity will be less reliable. This is due tothe increased dissociation between the electrical and mechanicalproperties of abnormal myocardium. Because of this, materialcharacteristics may be modified as to generate specific raw signalvoltage waveforms that are more characteristically seen with contractionor expansion. With such CMES characteristics, signal processing toderive the inferred polarity can be simplified and the resultinginformation more accurate.

In some embodiments as shown in FIG. 12, the apical and basal CMESsignals are processed at blocks 1200, 1202. The processed signals arethen vector labeled based on temporal and morphologic characteristics atblock 1220. At block 1240, a subtraction function is utilized tocalculate the difference in rotation between Aa 1112 and Ab 1114 (bottomof FIG. 11). The accuracy of the subtraction function is dependent uponappropriate vector labeling. At block 1260 a torsion calculator isoptionally implemented to generate data in numerical format that iscommunicated 1280 from the device to the programmer via wirelesstelemetry. Alternatively, some of the processes shown in FIG. 12 can beperformed within the programmer itself. In some embodiments, torsionalvelocity calculation is performed by analysis of relative values frombasal and apical CMES sensors as described above. Derivation of arotational displacement index can be performed by integration of thederived rotational velocity waveform.

It is noteworthy to mention that a combination of the forces generatedduring isovolumic contraction and relaxation will contribute to thedevelopment of the CMES signal and direction specific information maynot always be able to be characterized. Thus, in some embodiments, CMESdata can provide a crude representation of deformation and/or motion.The more myocardial surface area the CMES material covers, the morephysiologically accurate the derived indices will be at characterizingthe mechanical events occurring during isovolumic contraction andrelaxation. It is also noteworthy to mention that the temporalcharacteristics of the raw CMES voltage signal occur on or about thetime of mitral valve and aortic valve closure, but are only temporallyrelated to these events rather than representative of valvularmechanics. Under circumstances where the CMES sensor is free floating,myocardial acceleration (and possibly dP/dt, the rate of change in bloodpressure at the sensor site) and acoustical information may be derived.

Any and all of the data described herein can be used for monitoringcardiac performance and properties of dysynchrony. Likewise, the samedata can be implemented for optimization of interval timing for anymulti-site pacing system in a closed loop fashion as depicted in FIG. 9which describes the Matrix Optimization Method.

In an alternate embodiment, periodic interval monitoring is used toderive any of the indices described herein. During time frames wherediagnostic data is not collected, the voltage generated from the CMESsis stored as energy to reduce the costs to the system (e.g. batterylongevity) of operating such software.

FIG. 13 represents actual porcine animal data with an embodiment of aCMES in the ring or proximal position of a pacing electrode in the RVapex, where the electrode tip is in tissue contact but the CMES is notin close contact to myocardium. The top signal is a surface ECG 1300,the middle signal is a first order derivative, dCMES/dt 1302, similar inquality and in its temporal relationship to tissue Doppler derivedmyocardial velocity time graphs depicted and described above (FIG. 5B).The integral of this data, CMES 1304, bottom signal, is displacement.Peak systolic RV apical displacement is identified in the figure byarrow 1306. Comparable data can be acquired from a larger surface areaCMES that is basally located and parallel to the cardiac longitudinalaxis. A higher fidelity signal more representative of global cardiacdisplacement can be derived from such a sensor. Summation averaging ofmultiple waveforms will provide data with improved signal to noiseratio.

FIG. 14 represents actual porcine animal data with an embodiment of aCMES sensor in the LV anterior interventricular vein located ⅔ thedistance from the apex toward the base parallel to the cardiaclongitudinal axis. The sensor is in contact with the underlying tissue.The waveforms 1400 and 1402 derived are more representative ofmyocardial deformation and strain. Thus, dCMES/dt is an index of strainrate and the integral of this provides an index of strain. Thestrain/strain rate time graphs are similar to those acquired usingtissue Doppler imaging and speckle tracking techniques described above.Inversion of the waveforms 1400 and 1402 to derive an analogous vector1500 and 1502, as shown in FIG. 15, demonstrates waveforms similar tothose depicted in FIG. 5B derived from the strain rate equation (Eq. 2)being applied to Doppler derived myocardial tissue velocity imagingperformed, for example, by GE Vivid series echocardiography equipment.Arrow 1504 is peak longitudinal strain.

Second order derivatives of displacement data or first order derivativesof velocity data can be used to calculate acceleration indices as well.

Interval Specific Ensemble Averaging

As shown in FIG. 16, averaging of the measured mechanical stresswaveforms is synchronized with the detected heart events, such asspontaneous R-waves 1600 or stimulated events such as valvular heartsounds 1602 detected by an implanted sonomicrometer, filtered and/orprocessed CMES signal, or a signal from an alternate sensor.Synchronization of data acquisition can also be triggered by animpedance based parameter or index that relates to respiration and/ormyocardial mechanics. The waveform 1608 is averaged over a predeterminednumber of consecutive heart cycles 1606 by taking the sample average forevery time distance from the detected heart event, such as a QRS complex1604. The number of predetermined heart cycles could for instance be 30.For example, if the sampling frequency is 1 kHz, an averaged samplevalue at 24 ms distance from a QRS is calculated by taking the value at24 ms distance from a QRS for the predetermined number of heart cycles,which is 30 in this example, and then averaging the values. Theaveraging is repeated for all samples in the heart interval. This willresult in an averaged waveform 1610 based on the predetermined number ofbeats (in the example 30 beats). The strain calculations are thenperformed using the averaged waveform 1610. The advantage is that shortterm variations depending on for instance respiration are cancelled out.This method of averaging is referred to as “waveform averaging”. Havingthe advantage of enhancing details in the acquired waveform, the heartrate is preferably fairly stable during the process. This methodologycan improve signal-to-noise ratio. Data acquisition during periods ofrest and relative apnea or hypopnea will further improve the signal tonoise ratio (SNR). As shown in FIG. 17, input from a can basedaccelerometer 1700, to determine whether the patient is at rest, andrespirometers 1702, to determine whether the patient is in a state ofapnea or hypopnea, can trigger times for CMES data acquisition 1704 andfunction in conjunction with the Interval Specific Ensemble Averagingfeature describe herein and in FIG. 16.

Value Averaging

An alternative method to perform the averaging is to calculate thestrain parameters for each non-averaged consecutive heart beat and thenaverage the calculated parameters over the predetermined number of heartbeats. This method of averaging is referred to as “value averaging.”Having the advantage of detecting beat-to-beat variations of themeasured parameters, the heart rate does not have to be fairly stableduring the process. This is particularly suitable when variabilityanalysis is to be performed on the measured parameters.

Other Averaging Techniques

The average calculation above is performed using consecutive heartbeats, numbered 1, 2, 3 . . . , and so on. Alternatively, two averagevalues can be calculated. For example, the first value can be calculatedusing odd numbered beats: 1, 3, 5 . . . , and so on. The second valuecan be calculated using even numbered beats: 2, 4, 6 . . . , and so on.The two averaged values can be expected to be equal, but during severeheart tissue ischemia two groups can be formed. This will be the resultof the 2:1 rhythmic pattern in heart beats often seen during thiscondition. Other manifestations are the presence of rhythmic T-wavealternans (TWA) and pulsus alternans. Processing the measured strain inthis way forms a strong detector for this condition and can serve tonotify the clinician that a change in physiologic status has occurred.

1. An implantable cardiac stimulation device comprising: a first leadadapted to be implanted in or on the heart of a patient and to providetherapeutic stimulation to the heart of the patient, wherein the firstlead comprises a first mechanical sensor that obtains measurementsindicative of physical contraction and relaxation of the walls of theheart during systole and diastole; a controller that induces a deliveryof therapeutic stimulation to the heart of the patient via the firstlead, wherein the controller receives signals from the first mechanicalsensor indicative of contraction and relaxation of the walls of theheart, develops a template signal that corresponds to observedcontraction and relaxation of the walls of the heart during systole anddiastole, and uses the template signal to modify the delivery oftherapeutic stimulations being provided to the heart so that the heart'sfunction during systole and diastole is improved.
 2. The device of claim1 wherein the first mechanical sensor is adapted to be implantedadjacent a distributed region of a chamber of the heart substantiallyalong a first axis and wherein the controller processes the signals fromthe first mechanical sensor so as to develop a strain related parameter,the first mechanical sensor comprising a thin piezoelectric materialconfigured to contact a large amount of myocardial tissue surface area.3. The device of claim 2 wherein the controller processes the strainrelated parameter over time so as to develop one or more of a strainrate related index and a velocity index.
 4. The device of claim 1further comprising a second lead, wherein the second lead comprises asecond mechanical sensor that provides signals to the controllerindicative of contraction and relaxation of the walls of the heart. 5.The device of claim 4 wherein the first mechanical sensor is adapted tobe positioned at the interventricular septal region of the heart, andthe second mechanical sensor is adapted to be positioned in the lateralregion of the left ventricle.
 6. The device of claim 5 wherein thecontroller processes the signals from the first mechanical sensor andthe second mechanical sensor to develop a dysynchrony index.
 7. Thedevice of claim 6 wherein the dysynchrony index comprises a ratio of atime of peak signal from the first mechanical sensor to a time of peaksignal from the second mechanical sensor.
 8. The device of claim 6wherein the controller induces the delivery of therapeutic stimulationto induce greater synchrony based on the dysynchrony index.
 9. Thedevice of claim 8 wherein the therapeutic stimulation comprises changingone or more of atrioventricular timing and interventricular timing. 10.The device of claim 4 wherein the first mechanical sensor is adapted tobe positioned in the apical region of the heart and the secondmechanical sensor is adapted to be positioned in the basal region of theheart.
 11. The device of claim 10 wherein the controller process signalsfrom the first mechanical sensor and the second mechanical sensor todevelop one or more of a rotational index and a torsional index.
 12. Thedevice of claim 1 wherein the first mechanical sensor is adapted to bepositioned along a first axis that generally corresponds to thelongitudinal axis of a ventricle of the heart.
 13. The device of claim 1wherein the first mechanical sensor is adapted to be positioned in oneof the proximal to lateral portion of the main coronary sinus branch,the atrioventricular groove and the atrioventricular ring.
 14. Animplantable cardiomechanical assessment device comprising a controllerconfigured to: accept input relating to cardiomechanical strain of alateral region of a left ventricle of a heart; accept input relating tocardiomechanical strain of an interventricular septal region of theheart; and compute an interventricular dysynchrony index based upon thecardiomechanical strain input from the lateral region of a leftventricle of a heart.
 15. The device of claim 14 wherein the controlleris further configured to determine the times of peak cardiomechanicalstrain from the inputs.
 16. The device of claim 14 wherein the inputrelating to cardiomechanical strain comprises the output of acardiomechanical sensor.
 17. The device of claim 14 wherein theinterventricular dysynchrony index is used to set an interventricularinterval timing.
 18. A method for assessing myocardial function usingcardiomechanical sensors, comprising: acquiring data over a period oftime from a first implanted myocardial mechanical sensor and a secondimplanted myocardial mechanical sensor separated by a distance;summating the acquired data from the first and second implantedmyocardial mechanical sensors; and taking a derivative of the summatedacquired data over the period of time to determine a first strain rateindex.
 19. The method of claim 18 wherein the distance between the firstimplanted myocardial mechanical sensor and the second implantedmyocardial mechanical sensor is less than about 10 mm.
 20. The method ofclaim 18 further comprising multiplying the first strain rate index bythe distance between the first implanted myocardial mechanical sensorand the second implanted myocardial mechanical sensor to generate amyocardial velocity index.
 21. The method of claim 20 further comprisingintegrating the myocardial velocity index over the period of time. 22.The method of claim 18 wherein the first implanted myocardial mechanicalsensor and the second implanted myocardial mechanical sensor are both incontact with a patient's myocardium.
 23. The method of claim 18 whereinthe first implanted myocardial mechanical sensor and the secondimplanted myocardial mechanical sensor are not in contact with apatient's myocardium.
 24. The method of claim 18 wherein the firstimplanted myocardial mechanical sensor is contacting the patient'sinterventricular septum and the second implanted myocardial mechanicalsensor is contacting the patient's left ventricular lateral wall. 25.The method of claim 18 wherein the first implanted myocardial mechanicalsensor is implanted in or on a right ventricle of a patient and thesecond implanted myocardial mechanical sensor is implanted in or on aleft ventricle of the patient.
 26. The of claim 18 further comprising:correlating the acquired data over the period of time to a first portionof the patient's cardiac cycle; and averaging the acquired dataassociated with the repeating first portion of the patient's cardiaccycle.
 27. The method of claim 26, further comprising: correlating theacquired data over the period of time to a second portion of thepatient's cardiac cycle; and averaging the acquired data associated withthe second portion of the patient's cardiac cycle.
 28. An implantablecardiomechanical assessment system comprising: an implantablecardiomechanical sensor system comprising at least a first myocardialmechanical sensor and a second myocardial mechanical sensor; and animplantable controller system coupled to the implantablecardiomechanical sensor system and configured to acquire data over aperiod of time from the first implanted myocardial mechanical sensor andthe second implanted myocardial mechanical sensor, summate the acquireddata and calculate a derivative of the summated acquired data todetermine a first strain rate index.
 29. The system of claim 28 whereinthe implantable cardiomechanical sensor system comprises a first sensorlead and the first myocardial mechanical sensor and the secondmyocardial mechanical sensor are both located on the first sensor lead.30. The system of claim 29 wherein the implantable cardiomechanicalsensor system further comprises a second sensor lead, a third myocardialmechanical sensor and a fourth myocardial mechanical sensor, whereinboth the third and fourth myocardial mechanical sensors are located onthe second sensor lead.
 31. The system of claim 29 wherein the firstmyocardial mechanical sensor is oriented longitudinally with respect tothe first sensor lead.
 32. The system of claim 29 wherein the firstmyocardial mechanical sensor is oriented helically with respect to thefirst sensor lead.